Tissue engineering typically involves the synthesis of biologically relevant tissue for a wide range of applications including the replacement or support of damaged organs. A common strategy is culturing target specific cells in vitro in a scaffold followed by implantation of the scaffold into a biological organism. As a logical cellular source for tissue engineering, stem cells have attracted a great deal of attention due to their relatively fast rate of proliferation and diverse differentiation potential for various phenotypes. These include cells derived from several origins: induced pluripotent stem cells from fibroblasts, mesenchymal stem cells from bone marrow and adult stem cells from adipose tissue. Stem cells self-renew and their terminal differentiation depends on the influence of certain soluble molecules (e.g., growth factors, cytokines, etc.) as well as physical and biochemical interactions with the scaffold. Cellular behavior and subsequent tissue development at the cell-scaffold interface, therefore, involve adhesion, motility, proliferation, differentiation and functional maturity. The physicochemical properties of a scaffold, such as bulk chemistry, surface chemistry, topography, three-dimensionality and mechanical properties, all influence cellular response. Bulk chemistry can control cytotoxicity, as most scaffolds are made of biodegradable materials and must eventually release the by-products of their degradation. The effect of surface chemistry is often mediated by instantly adsorbed proteins such as fibronectin, collagen, fibrinogen, vitronectin, and immunoglobulin that affect phenotype, viability, and morphology, as well as proliferation and differentiation.
Studies regarding the effect of surface topography and texture on cellular response have been conducted. Stem cells are known to recognize topographical features on the order of hundreds of nanometers to several micrometers and exhibit distinctive genomic profiles in the absence of biochemical differentiation cues as well as a commitment to terminal differentiation. Electrospun scaffolds are ideal matrices for three-dimensional (3D) culture of cells and provide non-woven nano- to micro-sized fibrous microstructures typically having relative porosities of 70-90%. Natural biodegradable materials such as collagen, gelatin, elastin, chitosan, and hyaluronic acid, as well as synthetic biodegradable polymers such as poly(e-caprolactone) (PCL), poly(glycolic) acid (PGA) and poly(lactic) acid (PLA), have been electrospun for chondral and osseous applications.
In general, the broad utility of electrospun scaffolds for tissue engineering, wound healing, and organ replacement is clear (see Modulation of Embryonic Mesenchymal Progenitor Cell Differentiation via Control Over Pure Mechanical Modulus in Electrospun Nanofibers, Nama et al., Acta Biomaterialia 7, 1516-1524 (2011), which is incorporated by reference herein in its entirety, for all purposes) and the present invention provides, more specifically, polymer fiber constructs for use in the creation of nanofiber patches, as well as conduits for use in arteriovenous shunts for hemodialysis and blood vessel graft applications.
With regard to the creation of arteriovenous shunt grafts, the dysfunction of arteriovenous shunts in hemodialysis patients represents the single most common and burdensome complication in patients with end stage renal disease (see, US Renal Data System. USRDS Annual data report (2002); and Vascular Access in Hemodialysis: Issues, Management, and Emerging Concepts, Roy-Chaudhary et al., Cardiol Clin 23, 249-273 (2005), which are incorporated by reference herein in their entirety, for all purposes). More than 20% of all Medicare patients with end stage renal disease (ESRD) have vascular access graft complications that cost the U.S. healthcare system billions of dollars per year in access site treatment costs (see, Hemodialysis vascular access morbidity, Feldman HI, Kobrin S, Wasserstein A, J Am Soc Nephrol 7, 523-35 (1996) and Vascular Access in Hemodialysis: Issues, Management, and Emerging Concepts, Roy-Chaudhary et al., Cardiol Clin 23, 249-273 (2005), which are incorporated by reference herein in their entirety, for all purposes. In most of these patients, there are basically two approaches to establishing a vascular access site. The first approach is to create a native arteriovenous fistula by surgically anastomosing a larger artery to a vein, with the junction created by way of normal biologic healing of the anastomosis, eventually serving as a vascular access site for dialysis needle placement. However, due to underlying diseases in late stages, many patients present blood vessels that are non-viable for creation of a native AV fistula. In these patients, synthetic vascular grafts are implanted to create a shunt from the arterial side to the venous side of the circulatory system with the synthetic conduit serving as the location for vascular access. Expanded PTFE and cuffed double lumen silicone and urethane catheters are commonly used options for vascular access for hemodialysis with the latter being placed in a central venous site and the former typical implanted in the arm typically connecting a peripheral artery to a vein. E-PTFE grafts are preferred to central venous catheters because they perform better, although they are still far inferior to the preferred native fistulas.
The most common complication reported is low patency rates in ePTFE grafts, i.e., 50% at 1 year, and only 25% at 2 years (see Vascular Access in Hemodialysis: Issues, Management, and Emerging Concepts, Roy-Chaudhary et al., Cardiol Clin 23, 249-273 (2005) and FIG. 6, generally). Thus, only 1 in 4 patients with these synthetic grafts have unblocked (or patent) grafts at 2 years. The predominant source of this problem is that of neo-intimal hyperplasia, which occurs nearly 70% of the time at the venous anastomosis site of the graft, with the remainder of the graft blockages occurring at the arterial site or mid graft. Other complications include formation of pseudo-aneurysms in the graft (dilations), shredding of the graft wall due to repeated needle perforations, and thrombosis at needle perforations that do not close after needle removal. Secondarily, venous or arterial neo-intimal hyperplasia occurs due to the fact that the currently used ePTFE grafts are perceived by the host as a foreign body, eliciting an exuberant foreign body inflammatory response at the site of anastomosis, eventually resulting in hyperproliferation of cells that deposit undesirable tissue in the lumen of the graft or vein/artery at the anastomosis site or just downstream from the venous anastomosis location. In addition, current ePTFE conduits do not address compliance mismatch issues on the artery or venous side and clinicians point to this being another source of biological/biomechanical mismatch that results in hyperplastic response that leads to reduced patency rates. Another reason for lack of patency is the inability of ePTFE grafts to form a viable endothelial cell lining on the inner lumen, which if formed would be the ideal biological interface between the graft and the blood flow. Formation of a viable endothelial layer that presents the appropriate anti-thrombotic receptor site proteins to the blood flow would represent an ideal solution for AV shunt grafts. Finally, needle access sites in ePTFE grafts for dialysis are commonly reported sites for graft failures which can range from lack of closure/healing, blood stream and graft site infection, formation of pseudo aneurysms, and destruction of the graft wall due to multiple perforations, all of which require graft removal and replacement. These problems occur primarily due to two reasons: (1) inability of ePTFE conduits to self-close needle perforation site, and (2) inability of ePTFE conduits to produce biologic healing at needle perforation sites. The ability of the graft wall materials to self-close would represent a step forward in terms of reducing the damage caused by needle perforation and would limit the various complications. However, the inability of the synthetic ePTFE graft to go through a normal wound healing process, similar to what happens during vascular access at a venous or arterial site is a huge drawback. Since current grafts do not allow for effective cellular infiltration, angiogenesis, and tissue integration, they function primarily as a passive conduit and are susceptible to perforation related pin holing failures. If the graft wall is engineered to completely biointegrate, then there is an opportunity for the cellular, vascular, and tissue components that are present as a composite structure within the graft wall to respond with a biological healing response starting with hemostasis, inflammation, proliferation, and tissue remodeling towards a healed puncture site. Thus there is an ongoing need for a synthetic, implantable construct that overcomes these and other deficiencies of the prior art.